Devices, chambers, and apparati for culturing cells are a major focus in the biotechnology industry as systems for producing cells and cell-derived compounds as well as for using cells in tissue-engineering products as systems for therapeutic treatment and gene therapy. Although a wide variety of devices for cell culture have been conceived, developed, and applied in the last century, the need for novel and improved systems remains, in part because of unresolved limitations with existing devices and in part because of new applications with requirements not anticipated by existing devices. Unresolved limitations with current devices for cell culture include: (1) the inability to seed and distribute cells in devices at the relatively high ratios of volume of medium to cell number necessary for supporting cell inoculation (e.g., adhesion and spreading for attachment dependent cells) while reducing the volume per cell for advantageous cell function, growth, and concentration of cell-derived products, (2) inherent deficiencies in scaling of performance of cell cultures with increases in size of devices, and (3) incapabilities with economically-feasible scales of manufacturing that insure compliance with regulatory concerns. These problems are compounded by requirements for decreased limitations in mass transfer, the need to minimize holdup volume of perfused devices to limit hemodilution during in vivo treatment using cell cultures, and maximizing function of finicky cell cultures (e.g., stem cells that rapidly differentiate into undesired lineages).
Consideration of existing devices for cell culture and, in particular, for devices suitable for culture of fragile adhesion-dependent mammalian cells, illustrates the limitations and deficiencies that warrant addressing. For example, almost every device for cell culture previously disclosed lacks the ability to control the volume contained within the chamber housing the cells without changing the component comprising the walls of the chamber itself and/or compromising the sterility of the chamber and supported cell culture. The two types of devices providing for variable volume during cell culture without compromising sterility of the culture that have been described, moreover, present significant restrictions that hinder their application. Several forms of a chamber for cell culture based on a bag, which conceptually could allow variable volumes for cultures, have been described previously (e.g., U.S. Pat. Nos. 5,686,304 and 5,714,384), but the flexible walls present in these and other bags do not provide tight control of volumes, do not provide rigid surfaces for culture of adherent cells, nor present chambers with well-defined geometries for well-defined perfusions (e.g., uniform hydrodynamic shear stresses required for many adherent cells). Deformability of a wall of a chamber in general (e.g., as practiced by U.S. Pat. No. 6,152,163), leads to these inherent limitations. Alternatively, U.S. Pat. No. 5,707,868 describes the use of a piston-based design as a variable-volume chamber for cell culture. This type of design, similar in concept to other piston-based designs for biotechnological applications described in U.S. Pat. Nos. 5,143,847, 6,007,472, and 6,290,910, are cumbersome mechanically and not well-suited to large, planar cultures of adherent monolayers.
A review of previous designs of devices for cell culture supports the need for creation of an apparatus for the scalable culture of cells between substantially parallel, rigid flat plates in which a relatively large volume can be used to seed the cells and the holdup volume within the chamber itself reduced for perfusion without opening or otherwise disassembling the system to compromise its liquid-tightness and sterility. Such a device also should not require extensive handling or disassembly of the device between seeding and subsequent perfusion, such as by removal of a seeding well, and should improve on the normally labor-intensive process of cell culture while facilitating aseptic processing. A fully closed system in which cells are pumped directly into a chamber without direct exposure to the outside atmosphere, allowed to settle and attach, seeding medium removed, and perfusion of defined medium or plasma established with minimal disassembly or exposure would be compatible with these requirements. Further, these desired characteristics are even more critical for larger devices because the risk of contamination increases with size of device and reliable loading of cells becomes more difficult.
The development and application of devices allowing the culture of cells at high densities is of special importance to extracorporeal treatments for patients with diseased or otherwise failing organs. Such devices have applicability as therapies for other patients with islet failure (e.g., in diabetes), kidney failure, failure of endocrine organs (e.g., the adrenal glands), and impaired hematopoiesis (e.g., in cancer of the bone marrow).
The application of new forms of cell culture devices for treatment of individuals suffering from impaired liver function is a particularly pressing need. Over 43,000 Americans die each year from liver disease, making it the tenth leading disease-related cause of death in the US. When liver disease progresses to liver failure, the mortality is 80% unless a compatible donor organ is found. As with other organs, there is a critical shortage of donor livers. Over 12,000 patients currently are listed as transplant candidates, but fewer than half that number of donor livers become available each year. Treatment with a liver assist device (LAD) would decrease the mortality associated with liver failure by stabilizing patients so that they are suitable candidates for a transplant, by supporting them until a suitable donor liver becomes available, and/or by preventing deterioration to the point where a liver transplant is required. Improving the pre-operative health of these patients would also increase transplant success, thereby decreasing the frequency of retransplantation and easing the demand for donor organs.
In cases of sudden or hepatic failure, which often occur as a result of viral infection or toxicity, treatment with a LAD would eliminate the need for a transplant by supporting these individuals until their own livers regenerate. Liver transplantation is currently the most expensive organ transplant procedure. Successful development of a LAD would consequently provide major benefits to the US in reduced deaths and health-care costs.
Extracorporeal devices for temporary liver support have been investigated since the 1960s. Two strategies have been explored in the development of liver assist devices: (1) non-biological devices based on hemoperfusion on sorbents, hemodialysis across selectively-permeable membranes, and plasma exchange (Malchesky, “Non-biological liver support: historic overview,” Artif. Organs 18: 342-347, 1994); and (2) biological devices that incorporate cells or cellular components (Yarmush et al., “Assessment of artificial liver support technology,” Cell Trans. 1: 323-341, 1992).
Non-biological devices have shown only limited efficacy, confirming that synthetic materials cannot replace the range and level of complex metabolic functions normally performed by the liver. On the other hand, a biological LAD in which hepatocytes are seeded on the outer surface of hollow fibers and blood or plasma circulates through the lumen of these fibers was proposed almost 25 years ago by Wolf and colleagues (Wolf et al., “Bilirubin conjugation by an artificial liver composed of cultured cells and synthetic capillaries,” Trans. Amer. Soc. Artif Int. Organs 21: 16-23, 1975). It is desirable in such a LAD to provide the range of functions provided by hepatocytes in healthy livers, including clearance of protein catabolic products (e.g., hemoglobin from the turnover of red blood cells), detoxification of xenobiotics (compounds foreign to an organism), gluconeogenesis, homeostasis for lipids, minerals, vitamins, and cofactors, and regulation of blood composition (e.g., by secretion of carrier proteins like albumin and clotting factors).
Current designs for a biological LAD use the inverse of this concept today. Modern designs are often based on providing critical liver function by supporting high-density hepatocyte suspensions in hollow fibers, with circulation of blood or plasma outside the fibers. In this design, intermittent extracorporeal liver function is to be provided until the patient recovers through liver regeneration or until a transplant becomes available. However, the design based on hollow fibers is limited by several factors, including: a) inadequate mass transport, b) lack of scalability for sizing, c) lack of modularity for flexibility in design and assembly, d) poor control over distribution of cells, particularly during loading, e) inadequate support of hepatocytes during seeding, including limitations in volume of supporting medium, f) incompatibility with aseptic processing, g) constraints for void volume on the perfusion circuit for the device, and h) dynamics in mixing between device contents and patients' plasma due to constraints in the design of interface between device and patient. For example, it is desirable to perfuse ex vivo relatively large numbers of cells, up to 10% or more of the approximately 2-5×1011 hepatocytes in a healthy adult, at high densities (to minimize dilutional effects on plasma during treatment) and with significant differentiated function.
Hollow fibers have been chosen for LADs on the basis of ready availability rather than demonstrated ability to support hepatocyte function. Perfusion of high-density hepatocyte cultures in hollow fibers has shown a lack of convincing benefit due to, among other reasons, transport limitations that undermine their support of high-density cultures. Such limitations are particularly acute for oxygen, which is required for both basic metabolic function as well as for initial steps in detoxification. Perfusion of oxygenated plasma or medium through or around a network of hollow fibers fails to address this problem because these aqueous liquids are poor carriers for oxygen and the associated distances for transport are relatively large. Modifications to the core hollow-fiber design (e.g., the use of a woven network of three independent sets of capillaries providing integral oxygenation as disclosed in U.S. Pat. No. 5,516,691) significantly complicate fabrication and incompletely address underlying transport limitations. They also lack the ability to orient hepatocytes in a more organotypic laminar configuration.
In recent years several designs for devices for culture of liver cells that address a subset of eight critical factors limiting the performance of hollow fiber-based LADs have been described. U.S. Pat. No. 5,658,797 describes a device for treating hepatocytes cultured on plate-like, gas-permeable slides. However, this device has a complicated radial geometry and requires culture of these cells within a complicated and otherwise restricting sandwich between collagen gels. U.S. Pat. No. 6,228,607 describes improvements to the concepts introduced in U.S. Pat. No. 5,658,797 by change to a Cartesian geometry for flow; however, limitations in the configuration of the culture and requirement of a liquid-permeable membrane intervening between perfusate and cells complicate its application. International PCT Application Publication No. WO 00/78932 addresses the above limitations by describing modular devices in which hepatocytes are cultured on gas-permeable, liquid-impermeable films in direct contact with perfusate. No means for loading cells by perfusion as a closed system or for changing volume of the chamber for cell culture without compromising sterility are disclosed, however, in this latter application. All of the above disclosures also do not fully address the configuration of systems for interfacing an extracorporeal LAD with a patient in liver failure and in need of treatment.